Method and apparatus for advanced x-ray imaging systems

ABSTRACT

The present invention pertains to an apparatus and method for X-ray imaging a human patient. A vacuum bell bonded to an X-ray radiation-permeable window that can emit X-ray radiation from a plurality of spots located 1 cm from its edge, a collimator, and a detector are used. A ring of stationary X-ray sources can also be used with a stationary collimator and a rotating slot collimator and detector. An X-ray beam can be aligned in an X-ray system by establishing a position of the beam with respect to a moving collimator at a number of points in time, monitoring the velocity of the collimator, navigating the beam to a calculated position of a hole in the collimator, and correcting the alignment of the beam based on the location of the beam on the detector.

RELATED U.S. APPLICATION

This application is a continuation application claiming priority fromthe co-pending U.S. non-provisional patent application Ser. No.13/440,394, Attorney Docket Number TRT-7, entitled “METHOD AND APPARATUSFOR ADVANCED X-RAY IMAGING SYSTEMS,” with filing date Apr. 5, 2012. U.S.non-provisional patent application Ser. No. 13/440,394 claims priorityto U.S. provisional patent application Ser. No. 61/472,128, AttorneyDocket Number TRT-7, entitled “Computed Tomography Systems WithStationary Source,” with filing date Apr. 5, 2011, all of which arehereby incorporated by reference in its entirety.

FIELD OF THE INVENTION

The present invention pertains to x-ray imaging systems. The presentinvention also pertains to x-ray computed tomography systems.

BACKGROUND

Computed tomography (CT) is undeniably an important technology in modernmedicine. CT is used throughout screening, treatment, and follow-upexaminations; it is not uncommon for a patient to receive ten CT scansduring this time.

In light of the increasingly frequent use of CT for cancer imaging, highradiation dose from CT scans are concerning. There is strong scientificevidence that the elevated dose exposures as found in CT are leading toa significantly increased risk of cancer, especially in young patients.The National Academy of Sciences BEIR V (Biological Effects of IonizingRadiations) committee and the ICRP (International Commission onRadiological Protection) have reported that for a single acute radiationexposure in children the lifetime attributable cancer mortality risk isas high as 14% per Gy. This risk goes down with age, but is still in the2-8% range in the 30-50 age group. Importantly, recent reports on thebiological effects of radiation reaffirm the utility of the linearno-threshold model of radiation risk for solid cancers. Brenner et al.have shown that the average dose of a single CT scan leads to about 5cancer deaths in 10000 patients depending on the age of the patient andtype of exam performed.

The magnitude of the problem is worsening due to the large number of CTscans performed every year (62 million in the United States in 2006).While CT accounts for only 17% of imaging procedures using radiation, itdelivers 49% of the overall dose. The NCI (National Cancer Institute)website states: “CT is the largest contributor to medical exposure tothe U.S. population.”

CT is currently being investigated as a screening tool for early,asymptomatic cancer detection. A recent publication in the New EnglandJournal of Medicine by the International Early Lung Cancer ActionProgram Investigators concludes: “Annual spiral CT screening can detectlung cancer that is curable.” The NCI is currently conducting theNational Lung Cancer Screening Trial (NLST) with the objective toqualify CT and other X-ray modalities for lung-cancer screening. Onepotential outcome of this trial is the recommendation to screen a largepopulation for lung cancer with CT, thus increasing the radiationexposure to a large group of asymptomatic people. In fact, a recentpress release reports that NLST “found 20 percent fewer lung cancerdeaths among trial participants screened with low-dose helical CT.”Additionally, NCI is funding an R01 grant to develop aniterative-reconstruction algorithm for dose reduction in CT lung-cancerscreening. Thus, the prospect of CT use in cancer screening warrantsinvestigating ways to make CT more dose efficient.

Colorectal cancer, the fourth leading cause of cancer deaths worldwide,is largely preventable with appropriate screening. Optical colonoscopyis the method of choice. However, the nature of the exam has led to lowcompliance with the recommended screening intervals. Virtual colonoscopywith CT is emerging as an alternative that potentially could lead tomuch higher compliance rates. For example, for President Obama's yearlyhealth checkup, virtual colonoscopy was chosen over optical colonoscopy.Again, if virtual colonoscopy gains traction as a screening tool, alarge number of asymptomatic people will receive regular CT scans withhigh radiation exposure.

Another area of immense concern is pediatric cancer imaging. A recentstudy by Robbins evaluates the treatment protocols by the Children'sOncology Group that are typically used in the United States. The studyshows that throughout diagnosis, treatment, and follow-up periods forchildhood cancers such as neuroblastoma, Wilms tumor, Ewing sarcoma andlymphoblastic lymphoma, the radiation dose from imaging studies rangesbetween 109 and 152 mSv. The radiation dose is mostly from CT with theevaluated cases involving more than 15 CT scans each. Based on the BEIRV and ICRP reports, the lifetime risk of these children developing afatal cancer is in the 1-2% range. While pediatric cancer is a smallfraction in the overall cancer problem, children are a particularlyvulnerable group.

These examples highlight the broader problem of high radiation exposurein cancer imaging with CT. Therefore, it is desirable to reduceradiation dose in CT to continue the impressive success of CT infighting cancer and at the same time reduce the risk of causing cancerwith the very same modality.

Furthermore, modern computed tomography (CT) scanners have the goal ofcovering a large volume of the patient in a single rotation at very fastrotation speeds. This objective is driven by demands of cardiac CT tocover the entire organ in less than a heartbeat. Impressive results havebeen achieved with the current generation of CT scanners. However, thedownside of this development is the increased dose to the patient, theincrease in scatter, and the degradation of image quality in the outerslices due to cone beam artifacts. In particular, the increased dose inmedical imaging has come under scrutiny, with several published studiesdocumenting the elevated risk of cancer resulting from the radiationinvolved in medical imaging.

CT manufacturers are exploring a variety of methods to reduce this dosewhile maintaining image quality. However, these improvements areexpected to be minor compared to that which may be gained by analternative CT system concept, inverse-geometry CT (IGCT). Conventionalpoint source CT utilizes a single focal spot X-ray source and alarge-area detector, whereas IGCT utilizes a large-area, multi-focalspot X-ray source and a small-area detector. IGCT offers higher doseefficiency and faster acquisition times than state-of-the-artconventional point source CT systems. Thus, IGCT has the potential toovercome disadvantages with conventional point source CT andsignificantly out-perform conventional point source CT scanners.

However, IGCT as currently realized in prototypes faces difficulties inimplementation due to a large source array to be rotated at high speedsand significant challenges from high power and cooling requirements ofthe source.

What is needed is a CT imaging system capable of producing rapid highquality images. Furthermore, the CT imaging system should provide lowradiation imaging.

SUMMARY

In one embodiment, an X-ray system for imaging a human patient isprovided. A vacuum bell bonded to an X-ray radiation-permeable windowthat can emit X-ray radiation from a plurality of spots located 1 cmfrom its edge, a collimator, and detector are used. Theradiation-permeable window and vacuum bell can be bonded with a brazedor electron beam-welded connection. The radiation-permeable window maybe beryllium and may have a thin film tungsten target deposited on it. Asecond vacuum bell and radiation-permeable window can be used, and thefirst and second permeable windows can be in contact with each other.

In another embodiment, a ring of stationary X-ray sources with astationary collimator and rotating slot collimator and detector areprovided. The rotating slot collimator can span an arc between 60 and160 degrees. The stationary collimator can have at least ten or between10 and 50 slots, can have at least 10 slots perpendicular to the slotsof the rotating collimator, and can comprise metal rings. The system canuse cooling water to remove heat from the X-ray target and can beoperated at full power for at least one hour. The system can use asensor to monitor the velocity of the rotating collimator.

In another embodiment, an X-ray beam is aligned in an X-ray system byestablishing a position of the beam with respect to a moving collimatorat a number of points in time, monitoring the velocity of thecollimator, navigating the beam to a calculated position of a hole inthe collimator, and correcting the alignment of the beam based on thelocation of the beam on the detector. The position of the detector canbe calculated using the initial position and velocity of the collimator.The centroid position of the X-ray beam on the detector can bedetermined and compared to its calculated value. The X-ray beam can beaborted if the X-ray beam is not aligned.

These and other objects and advantages of the various embodiments of thepresent invention will be recognized by those of ordinary skill in theart after reading the following detailed description of the embodimentsthat are illustrated in the various drawing figures.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention is illustrated by way of example, and not by wayof limitation, in the figures of the accompanying drawings and in whichlike reference numerals refer to similar elements.

FIG. 1 is a diagram showing an exemplary fixed-source computedtomography imaging system of one embodiment of the present inventionwith three vacuum envelopes.

FIG. 2 is a diagram showing an exemplary isolated collimator-detectorassembly of one embodiment of the present invention.

FIG. 3 is a diagram showing an exemplary fixed-source computedtomography imaging system of one embodiment of the present inventionwith nine vacuum envelopes.

FIG. 4 is a diagram showing an exemplary uniform illumination pattern.

FIG. 5 is a diagram showing an exemplary illumination pattern thatprovides an increased flux in the central region of the collimator.

FIG. 6 is a diagram showing an exemplary illumination pattern thatprovides an increased flux in the central region of the collimator.

FIG. 7 is a diagram showing an exemplary source ring with linear sourcesof one embodiment of the present invention.

FIG. 8 is a diagram showing an exemplary arrangement of sources of oneembodiment of the present invention.

FIG. 9 is a diagram illustrating a stationary collimator ring androtating collimator arc of one embodiment of the present invention.

FIG. 10 is a diagram illustrating a collimator alignment configurationof one embodiment of the present invention.

FIG. 11 is a diagram illustrating a signal from a detector with correctcalculated collimator position in one embodiment of the presentinvention.

FIG. 12 is a diagram illustrating a signal from a detector withincorrect calculated collimator position in one embodiment of thepresent invention.

FIG. 13 is a diagram illustrating the illumination of two detectors bytwo closely abutted X-ray sources in one embodiment of the presentinvention.

FIG. 14 is a diagram illustrating components of a scanning-beam digitalX-ray source of one embodiment of the present invention.

DETAILED DESCRIPTION

Reference will now be made in detail to embodiments of the presentinvention, examples of which are illustrated in the accompanyingdrawings. While the invention will be described in conjunction withthese embodiments, it will be understood that they are not intended tolimit the invention to these embodiments. On the contrary, the inventionis intended to cover alternatives, modifications and equivalents, whichmay be included within the spirit and scope of the invention as definedby the appended claims. Furthermore, in the following detaileddescription of embodiments of the present invention, numerous specificdetails are set forth in order to provide a thorough understanding ofthe present invention. However, it will be recognized by one of ordinaryskill in the art that the present invention may be practiced withoutthese specific details. In other instances, well-known methods,procedures, components, and circuits have not been described in detailas not to unnecessarily obscure aspects of the embodiments of thepresent invention.

FIG. 1 is a diagram showing an exemplary fixed-source computedtomography imaging system of one embodiment of the present invention.Imaging system 100 comprises a ring of X-ray sources 101, 102, and 103with an inner diameter of 1 m. The source ring can be made of threeX-ray sources 101, 102, and 103 making a three-gap system as shown. Forthe three-gap system, there can be three separate source arrays, eachcontaining three electron guns in a single vacuum envelope. Each ofthese source arrays can have a large-area tungsten transmission target.The source-spot locations can cover the full 360 degrees, except for asmall gap of a few centimeters between each of these arrays. The axialextent of the source array can be 16 cm. There can be a fixedpre-collimator between the source arrays and the spinning ring. Thispre-collimator defines the possible locations of the source-spots.

Within the ring of X-ray sources 101, 102, and 103 can be a rotatingdetector/collimator assembly. In one embodiment, only detector 110 andcollimator 120 rotate. Collimator 120 can consist of an array of holeswith each hole capable of illuminating the entire detector array. Thecenter of the detector array can be diametrically across from the centerof the collimator arc.

In operation, each row of the collimator 120 can have X-rays firingthrough its holes starting, for example, with the trailing hole andmoving sequentially to the leading hole. The collimator rows can fire insequence. A “super-view” can be obtained after all holes of allcollimator rows have “fired”. Other firing sequences are possible.

The detector elements can be read after a source-spot fires. The axialwidth of the detector array can also be 16 cm. By using the same axialwidth for both source and detector arrays there are no rays outside ofthe region of interest in the axial direction. Thus there is no unusedexposure such as occurs in cone-beam systems.

Imaging system 100 can have a large, 100 cm diameter, stationary ring ofscanning X-ray source-spots. Inside the source-ring can be a rotatingring containing detector 110 and collimator 120. This rotating ring, organtry, obtains power and outputs the detector signals through aslip-ring. FIG. 2 is a diagram showing an exemplary isolatedcollimator-detector assembly of one embodiment of the present invention.Collimator 220, which can be mounted opposite detector 110, can have ahole-pattern that focuses the X-rays onto detector 110. Detector 110 canbe 6 cm by 16 cm and collimator 220 can span an arc of about 120 degreesand can have a width of 16 cm. Each collimator holes can illuminate theentire detector. This system design allows for rotation speeds of atleast three rotations per second with image quality comparable to aconventional point source CT scanner.

Detector 110 can be a 6 cm by 16 cm detector. The detector ASIC can bemodified to allow parallel current-integration readout and dual-energyacquisition.

Collimator 220 can be designed to attenuate 120 keV photons. It canconsist of nearly 9,000 holes with a hole pitch of 2.3 mm. Each hole canbe tapered and angled to project X-rays onto a 5-cm by 10-cm detector ata distance of 150 cm. Collimator 220 can also have approximately 600holes projecting onto a 6-cm by 16-cm detector at a distance of 100 cm.Collimator 220 can be curved and have a larger area.

X-ray sources 101, 102, and 103 can be designed for continuous operationat 25 kW and at a tube voltage that can vary between 70 KVp and 120 kVp.The focal spot size can be 0.4 mm and the spot dwell time can be 1 μswith a duty cycle of 80%. The complete collimator can be scanned every15 ms. X-ray sources 101, 102, and 103 can include a thin-film tungstentarget layer deposited on a water-cooled 25-cm-diameter beryllium disc.The source power can be increased to 50 kW.

For a three-gap system, X-ray sources 101, 102, and 103 must cover asignificantly larger target area. A large vacuum envelope that housesthree guns in each source can be used. Each gun can illuminate a thirdof the target area. The use of three guns enables the entire target areato be illuminated. Different window material such as stainless steel andaluminum nitride can be used.

The projection data can be acquired as the collimator-detector assemblyrotates around the patient. Collimator 120 can be located between thesource array and the patient and source-spots are active only whenbehind collimator 120. Collimator 120 moves only a small angularincrement during the time the scan of every designated hole incollimator 120 is completed. A complete scan of collimator 120 isdescribed as a “superview”. The maximum travel of detector 110 during anacquisition of a superview is one detector width. Therefore, a completedataset can be obtained with as few as about 60 superviews.

High-weight, high-voltage, and high-power components of imaging system100 can be removed from the challenging environment of the rotatinggantry. Miniaturization of the high-voltage power supply is notrequired. High-power slip rings are not required. The X-ray source arraycan be cooled with hospital water, eliminating the conventionalgantry-mounted radiator and increased air-conditioning requirement.Faster rotation times and faster volume acquisitions are possible.Overall reliability can be increased by the removal of many components,especially X-ray sources, from the high-G-force environment of therotating gantry. A total source array area that is approximately threetimes larger than conventional point source systems can be required.However, the engineering necessary for this is greatly simplifiedcompared to a rotating source array. Also, the cost per area for sourcesis significantly less than the cost per area of detectors. Thus, thiscan be also economically feasible.

FIG. 3 is a diagram showing an exemplary fixed-source computedtomography imaging system of one embodiment of the present inventionwith nine vacuum envelopes. In this embodiment, the source ring is madefrom nine individual X-ray sources forming a nine-gap system. There arenine separate source arrays, each containing a single electron gun in asingle vacuum envelope. Each of the envelopes can have a 400 cm² sourcearea.

FIG. 4 is a diagram showing an exemplary uniform illumination pattern.FIG. 5 is a diagram showing an exemplary illumination pattern thatprovides an increased flux in the central region of the collimator. FIG.6 is a diagram showing an exemplary illumination pattern that providesan increased flux in the central region of the collimator. FIG. 5 andFIG. 6 show patterns with increased intensity in the center. In practicethe uniform and non-uniform patterns could be interleaved to ensuresampling completeness. The use of different illumination patterns canprovide a two-dimensional adaptive filter. Several schemes for selectingthe distribution of illumination are possible. Importantly, theillumination for one superview can be based on the results of theprevious superview.

Iterative reconstruction methods can also be used. In particular,Maximum Likelihood Expectation Maximization (MLEM) is well suited fordatasets from unconventional geometries. The algorithm is less prone tounder-sampling artifacts and tends to reduce noise compared to standardalgorithms.

One of the most critical design issues is the ability to produce enoughphotons to provide the desired image quality. Imaging system 100acquires enough photons to produce an acceptable image. The detectorarray is 60 mm by 160 mm giving an area of about 96 cm². The duty cycle(the source-spot on time) utilized of imaging system 100 can be 80%. Thesource of imaging system 100 can have a power rating of 50 kW. Comparedto a 85 kW tube, this reduces the number of photons by a factor of 0.59.Imaging system 100 can have a slightly shorter focus-to-detectordistance giving it a factor of 1.17 advantage.

Imaging system 100 does not rely on the anti-scatter grids used inconventional CT systems to reduce scattered radiation in the projectionimages. As discussed earlier, imaging system 100 takes advantage of thesignificantly smaller detector compared to a conventional system.Scatter scales approximately with the detector size assuming a constantdistance between patient and detector. The smaller detector of imagingsystem 100 can be a significant advantage as the amount of scatterscales with the illuminated volume that, for a fixed object, isproportional to the detector area. The amount of scatter can be lessthan 10% for imaging system 100 while for a conventional system scatterexceeds 40% In a conventional system, scatter is managed with ananti-scatter grid, whereas in imaging system 100, an anti-scatter gridwill not be necessary. The efficiency is about 75%. Imaging system 100can have a significantly lower scatter fraction and can be operatedwithout an anti-scatter grid, giving a photon advantage of 1.33.

Detector 110 can be photon counting, having an intrinsic DQE advantageof 20%. Additionally, photon counting detectors have a bias towardslower energies giving another 20% advantage. Thus, fewer photons areneeded for the same image quality and can be counted as a (virtual) fluxincrease of a factor 1.44.

The transmission anode of imaging system 100 can provide 1.7 times asmany photons for the same current as the more traditional steep-anglereflection anode.

Because imaging system 100 can adjust the number of photons dependingupon the thickness of the object on a view-by-view, or evenbeam-by-beam, basis, a significant increase in maximum number of photonscan be obtained. An average increase of a factor of 4 can be achieved.

The following table summarizes the cumulative advantages anddisadvantages, and shows that the number of available photons iscomparable to that of a standard point source system.

IGCT property relative to IGCT/Standard Cumulative standard point source0.15 0.15 smaller detector area 0.80 0.12 lower duty cycle for IGCT 0.590.07 less tube power 1.17 0.08 shorter source-detector distance 1.330.11 operation without AS grid 1.44 0.16 photon-counting detector 1.700.27 transmission anode 4.00 1.08 virtual bow-tie

The duty cycle can be increased to 100%. Imaging system 100 can usemultiple tubes that can be alternated thus filling in the off-time of asingle source. In addition, both iterative reconstruction and energyresolving detectors can improve performance. Overall, imaging system 100can increase the effective number of photons by more than a factor oftwo.

Some of the effects discussed previously convert directly into dosesavings to the patient. Imaging system 100 does not rely on theanti-scatter grids used in conventional CT systems. Anti-scatter gridsare positioned after the patient and also prevent a significantpercentage of the radiation from reaching the detector. Thus removingthe anti-scatter grid reduces the dose to the patient. The omission ofanti-scatter grids, and similarly, the removal of the dead-space betweendetector elements, leads to about a 25% improved dose efficiency.

The implementation of an adaptive filter can be used with inversegeometry CT and imaging system 100. The effective intensity of eachsource-spot-to-detector beam can be adjusted depending on the patientthickness, or attenuation, for that beam. This adaptive approach alsominimizes irradiation where no body parts are present. A dose saving onthe order of a factor of two can be achieved. Photon counting detectorsprovide an additional dose savings of a factor of 1.44.

The combined dose saving with imaging system 100 is almost a factor of4. Even further dose savings can be achieved with the use of an energyresolving detector and iterative reconstruction methods. Imaging system100 can be used only to scan the organ of interest and thereby furtherreduce the dose to the patient.

FIG. 7 is a diagram showing an exemplary source ring with linear sourcesof one embodiment of the present invention. Rather than using atwo-dimensional array of sources, the source ring uses lines of sources.These linear sources can be constructed using either transmissiontargets or reflection targets. An array of linear X-ray tubes isarranged in a ring. Detector-collimator assembly rotates inside thering.

FIG. 8 is a diagram showing an exemplary arrangement of sources of oneembodiment of the present invention. An array of linear x-ray tubes isarranged in a ring. Detector-collimator assembly rotates inside thering. This arrangement of source-spots can achieve complete sampling asthe gap between any two linear sources is covered by a third linearsource as shown in FIG. 8. Every plane intersecting the ring alsointersects a source trajectory. As an example, although the dashed linelies in the gap between tube 2 and 3, it intersects tube 4 of FIG. 8.Another advantage is that the tube target, whether transmission orreflection, can be at a steep angle with respect to the X-ray beam. Thisallows a line-focus electron beam to be used which, in turn, enables afour-fold increase in tube power. The area source approach hasadvantages with the heat loading of the target and that implementationof the virtual bowtie is easier.

Imaging system 100 can have numerous advantages compared to conventionalpoint source CT systems. Imaging system 100 can have lower dose and canbe four-fold more dose-efficient than conventional point source systems.Imaging system 100 can have faster volume acquisition with scan timesless than 300 msec. Imaging system 100 can perform whole-organ imagingwith no table translation and no cone-beam artifacts. Data can bereconstructed using existing algorithms. Thus, advantages include fastacquisition and the reduction of dose, artifacts, and cost. Imagequality can be comparable to standard point source CT and also have asignificant margin to exceed current performance. Complete datasets canbe produced and a variety of reconstruction algorithms can be used forefficient reconstruction.

A new type of scanning beam X-ray tube is disclosed that will enableinverse geometry CT with the promise of reducing radiation dose by atleast a factor of four at comparable image quality as modern CTscanners.

Patient studies have shown that dose savings of a factor of 5 arepossible with a system of an embodiment of the present invention. Inaddition, an adaptive exposure technique whereby the X-ray exposure ismodulated depending on the local opacity of the patient can be used.Additional dose savings of more than 40% are possible; thus, potentiallythe system can operate at 9-fold lower dose than conventional systems.

Key to the advantages of inverse geometry imaging techniques are theavailability of a large number (thousands) of X-ray focal spotsdistributed over a large area. Such an imaging geometry allows thereduction of the solid angle of the X-ray beam. Rather than projectingonto a large detector with a single X-ray source as is done inconventional imaging geometries, in inverse geometries the image isformed by many projections from the different X-ray focal spots onto asmall detector. Hence, the solid angle of the X-ray beam issignificantly reduced. The advantage is that scatter on the detector issignificantly reduced as the scatter scales with the solid angle of theX-ray beam. Reduction in scatter reduces excess noise in the images andthus translates directly into dose reduction. Scatter reduction ininverse-geometry CT can translate into a dose saving of more than 40%.

Another advantage of small solid angles is that every projection imageonly samples a very small region of the imaging volume, thereforeallowing for adaptive exposure. In other words, rather than exposingevery part of the image with the same radiation dose, the exposure canvary depending on the opacity of the region exposed. For example,exposure can be reduced significantly in the lung field or thinner partsof the body and maintain exposure in more opaque regions. The dosesavings potential of adaptive exposure in CT is even higher than influorosocopy and can achieve dose savings of more than 50%.

The small active area of the detector allows for cost efficientimplementation of state-of-the-art detector technology. For example,photon-counting detectors have intrinsically better noise performance atlow photon counts and a beneficial bias towards lower energies ascompared to the energy-integrating detectors currently used inconventional CT systems. This can translate into dose savings of 20%.

An overall dose saving of a factor of 4 in inverse geometry CT (IGCT)can be achieved.

Earlier inverse-geometry CT designs envisioned rotating a large-areaX-ray source array and a small detector at high speeds similar to 3rdgeneration CT scanners where a large detector and an X-ray source arerotated. Disadvantages of this approach are the need to rotate largemasses at high speed, to have a high-power connection through a slipring, and no access to cooling water at the X-ray source.

One embodiment of the present invention comprises a stationary ring ofmulti-focus X-ray sources with a stationary slot collimator. The ring ofX-ray tubes can consist of a single vacuum envelope with one to 15electron guns or of 1 to 15 separate vacuum envelops with individualelectron guns or any combination of the former. Inside this ring, anannular sub-assembly can rotate, possibly consisting of a collimator onone side and a detector on the opposite side. The rotating collimatorcan have slots perpendicular to the slots of the stationary collimatorensuring collimation of the X-rays onto the detector. Advantages are thesignificant reduction of weight of the rotating assembly, elimination ofhigh-power connections over slip rings and the availability of coolingwater at the X-ray source.

Additionally, the target surface of the X-ray tube can be curved toconform to the collimator ring. A curved target with minimal inactivesource area between adjacent X-ray sources can be used. These inactiveareas can be chevron-shaped to minimize undersampling artifacts. TheX-ray source can consist of a cathode assembly and a target assembly.The X-ray source can use a cathode assembly that generates, accelerates,and deflects the electron beam. The target assembly can consist of aberyllium window and a vacuum bell. The window may also be made of amaterial other than beryllium, such as thin stainless steel, titanium,carbon, any combination thereof, or any other material with asufficiently low attenuation coefficient for X-rays to permeate and thatcan maintain the vacuum created within the bell. Alloys or combinationsof the above materials or other materials can also be used. The vacuumbell creates the vacuum envelope between cathode assembly and window. Astand-alone tube with circular target and significant inactive area inthe outer perimeter can be used. A tube that includes the chevron-shapedgap discussed earlier can also be used. Either a flat target, or acurved target that allows the use of a full ring of X-ray sources can beused.

Another engineering challenge can be the interface between the X-raytarget and vacuum bell. Currently, this is done using a stainless-steelring that is diffusion-bonded to the X-ray target assembly, which may,for example, consist of a thin-film tungsten target deposited on aberyllium window. This creates a relatively large inactive annular areaaround the X-ray target. The stainless steel ring is then welded to thevacuum bell. In one embodiment of the present invention, the window canbe bonded directly to the vacuum bell to eliminate the inactive area ofthe stainless-steel ring. Electron-beam welding can be used withberyllium. It may also be used with a number of other potential windowmaterials. Any other type of welding or bonding method may be used.

In other embodiments of the present invention, target materials includebut are not limited to tungsten, copper, molybdenum, and alloyscomprising these and/or other elements. A target may be deposited on awindow, which as previously discussed may be beryllium, thin stainlesssteel, titanium, carbon, any combination thereof, or any other materialwith a sufficiently low attenuation coefficient for X-rays to permeate,e.g. a material comprising element(s) with atomic number(s) less than30. The vacuum bell may stainless steel, copper, or other metals oralloys. A bonding method may be used which can attach the window to thevacuum bell directly. Depending on the materials of these twocomponents, electron beam welding, brazing, or other methods may beused.

The target can be utilized up to 1 cm from the outer edge of the X-raysource. Beyond IGCT, sources that can be closely abutted can be used inother medical imaging applications such as image guidance,tomosynthesis, or radiation-therapy monitoring. A curved target withminimal inactive source area between adjacent X-ray sources can be used.A flat target with minimal inactive source area between adjacent X-raysources can also be used. These inactive areas can be chevron shaped tominimize undersampling artifacts. Inactive areas can also be any othershape including but not limited to straight, curved, slanted, and soforth. The shape of the window, e.g. a beryllium window can match theshape of the target. For example, the shape of a beryllium window can bea curved window with chevron-shaped ends. Brazing and electron beamwelding can be used as possible bonding methods. The vacuum bell and itsoverall shape can conform to the beryllium window shape. The X-ray tubecan allow for accessibility of X-ray focal spots as close as 1 cm fromthe physical edge of the source.

Experiments can be performed with a pinhole that can be placed atdifferent locations on the target surface. The electron beam can besteered on the transmission-target surface to illuminate the pinhole. Asa baseline, the pinhole can be placed at the center of the target. Thebeam profile can be measured as projected onto the detector. Thealuminum half-value layer of the beam and the fluence can be measuredand compared to the established values. The pinhole can then be placedas close as 1 cm from the edge of the tube. The beam shape can bemeasured and the focus coils can be used to adjust the beam shape tomatch the profile in the center. The half-value layer of the beam andfluence can then be measured. If these values deviate from the values inthe center, the pinhole can be moved further away from the edge of thetube and measurements repeated. This technique can establish how closeto the edge spots can be accessed.

The scanning-beam digital X-ray system differs significantly from thedesign of a conventional fluoroscope. In contrast to a conventionalfluoroscope, in which the X-ray tube has only a single focal spot, thescanning-beam digital X-ray tube is extended and consists of a scanningelectron beam dwelling sequentially at up to 9,000 focal-spot positions.The X-ray tube technology is very similar to a CRT tube, but rather thanusing a phosphor target generating visible light from low-energyelectrons, the scanning-beam digital X-ray tube uses a tungstentransmission target that generates X-ray photons from high-energyelectrons. Other X-ray emissive targets including but not limited tomolybdenum and copper can also be used. At each focal-spot position,X-ray photons are emitted towards the detector by use of a focusingcollimator, thereby projecting a small view of the imaging volume.

In contrast to a conventional fluoroscope in which the detector is largeand close to the patient, the scanning-beam digital X-ray detector issmall and far away from the patient. The conventional fluoroscope usesan energy-integrating detector, while the scanning-beam digital X-raysystem uses a photon-counting detector. These aspects can contribute tosignificant reduction in X-ray dose.

In conventional systems, the scatter fraction can be high and thisproblem can be addressed by the use of anti-scatter grids, whichunfortunately can also reduce the primary radiation and thereby reducedose efficiency. Even with anti-scatter grids, the detected scatterfraction in a conventional system can be as high as 56% for largepatients. These scattered photons can contribute to the noise floor anddegrade the contrast-to-noise ratios of images acquired by the system.In the scanning-beam digital X-ray system, the detector can be small andfar away from the patient. Therefore, the scattered photons can have amuch lower probability of reaching the detector. For the same largepatient as described previously, the scanning-beam digital X-ray systemcan have a detected scatter fraction of less than 10%. The small scatterfraction in the scanning-beam digital X-ray system can allow foroperation without an anti-scatter grid.

A photon-counting detector possesses a reduced noise floor that can beespecially useful in imaging situations with low flux. Additionally,photon-counting detectors do not have a photon-energy bias as can beseen with energy-integrating detectors. In an energy-integratingdetector, a 120-keV photon can produce twice as much signal as a 60-keVphoton. In a photon-counting detector, both photons contribute equallyto the signal. This enables lower-energy photons, which can generatehigh image contrast, to contribute more to the image than in aconventional system.

The peak voltage implemented by a conventional, e.g. regular pointsource geometry, fluoroscope system can range from 73 to 116 kVp inselected projections. Scanning-beam digital X-ray patient entranceexposures range from a reduction of 5-fold to a reduction of 10-foldrelative to such systems. In general, scanning-beam digital X-raycontrast-to-noise compare favorably for exposures about 7-fold lowerthan that of the conventional point source exposure numbers, roughly inagreement with the 8-fold reduction observed in phantom studies. Theseexposure savings are equivalent to a 3.5-fold effective dose saving inpatients.

Adaptive exposure—a new technique to save dose—can be implemented.Rather than exposing every part of the image with the same radiationdose, scanning-beam digital X-ray technology allows variation of theexposure depending on the opacity of the region exposed. Thus, exposurecan be reduced in translucent areas such as the lung field and exposurecan be maintained in more opaque regions.

A comparison of two images can be obtained from an anthropomorphicphantom mimicking a 90-kg male with and without adaptive exposure. Theimage with adaptive exposure can be imaged with 50% fewer photons thanthe image without adaptive exposure. Besides reducing the dose, adaptiveexposure can also effectively compress the dynamic range of images andthereby improve image quality. With implementation of adaptive exposurethe scanning-beam digital X-ray system can potentially produceequivalent image quality as conventional systems at 7-fold lower dose.Similar dose savings using inverse geometry in CT can be achieved.

In early proposed configurations of IGCT, the source array and detectorarray both rotated on a gantry.

An IGCT design with a stationary X-ray source array and a rotatingdetector can have powerful advantages. One advantage of such an IGCTsystem geometry is that all high-weight, high-voltage, and high-powercomponents can be removed from the challenging environment of therotating gantry. Thus miniaturization of the high-voltage power supplyis not required; high-power slip rings are not required; the X-raysource array can be cooled with hospital water, eliminating theconventional gantry-mounted radiator and increased air-conditioningrequirement; faster rotation times and faster volume acquisitions arepossible; and overall reliability can be increased by the removal ofmany components, especially the X-ray source, from the high-G-forceenvironment of the rotating gantry.

This IGCT design may require a total source array area that isapproximately three times larger than previously proposed IGCT systemswherein both source and detector rotate.

An extended X-ray source array can be used with an inverse-geometry CTsystem with a large, e.g. 60 to 140 cm diameter, stationary ring ofscanning X-ray source-spots. For some applications the source ringdiameter may also be less than 60 cm or greater than 140 cm. However, insome embodiments of the present invention, the source ring diameter maybe between 60 and 80 cm, 80 and 100 cm, 100 and 120 cm, or 120 and 140cm, inclusive, or any other integer or non-integer number of centimeterswithin the enumerate ranges (e.g. 60, 65, 70, 75, 80, 85, 90, 95, 100,105, 110, 115, 120, 125, 130, 135 or 140 cm) or any range between 60 and140 cm.

As will be discussed later, the stationary ring of scanning X-raysource-spots can also be partially collimated by a stationary slotcollimator. Inside the source-ring can be a rotating ring containing adetector and a collimator. This rotating ring, or gantry, can obtainpower and output the detector signals through a slip-ring assembly. Thecollimator, which can be mounted on the rotating ring opposite thedetector, can have a slot-pattern that focuses the X-rays onto thedetector. The detector can without limitation be 4 to 14 cm by 8 to 24cm, and the collimator can span an arc of about 60 to 160 degrees andcan have a width of 8 to 24 cm. This system design can allow forrotation speeds of at least 0.5 rotations per second (rps) with imagequality comparable to a conventional modern CT scanner. However, otherdimensions and rotation speeds can also be used. The detector can besquare, rectangular, trapezoidal, or any other shape. For example, thedetector may be 10 cm by 10 cm but may also be 10 cm by 20 cm, or haveany other square or rectangular dimensions, including but not limited to4 to 14 cm by 8 to 24 cm. The collimator can focus radiation onto thedetector, and may be curved to conform to the stationary collimatorring. It may span between 60 and 80 degrees, 80 and 100 degrees, 100 and120 degrees, 120 and 140 degrees, or 140 and 160 degrees, inclusive, orany other number of degrees within the enumerate ranges (e.g. 60, 65,70, 75, 80, 85, 90, 95, 100, 105, 110, 115, 120, 125, 130, 135, 140,145, 150, 155 or 160 degrees) or any range between 60 and 160 degrees.Alternatively, for some applications, the collimator may span less than60 degrees or more than 160 degrees.

Projection data can be acquired as the collimator-detector assemblyrotates around the patient. The collimator can be located between thesource array and the patient and source-spots can be active only whenbehind the collimator. The collimator may move only a small angularincrement during the time the scan of every designated hole in thecollimator is completed. A complete scan of the collimator may bedescribed as a “superview”. The maximum travel of the detector during anacquisition of a superview may be one detector width. A complete datasetcan be obtained with as few as about 60 superviews.

As in fluoroscopy, the dose savings can be achieved by scatterreduction, superior detector technology and adaptive exposure.

Scatter reduction may be significant in inverse geometry systems.Inverse geometry CT does not rely on the anti-scatter grids used inconventional CT systems to reduce scattered radiation in the projectionimages. As discussed earlier, IGCT takes advantage of the significantlysmaller detector compared to a conventional system. Scatter scalesapproximately with the detector size assuming a constant distancebetween patient and detector. Thus, in the IGCT system scatter impingingonto the detector is about 8% of that of a conventional system. In aconventional system scatter can be managed with an anti-scatter grid; inthe IGCT system an anti-scatter grid may not be necessary.

Operation without an anti-scatter grid can translate into significantdose savings as anti-scatter grids can have a reduced transmissivity andblock a large fraction of the primary X-rays. Transmissivity can beexpressed in terms of the primary transmission factor t_(p) of ananti-scatter grid. The ability of the anti-scatter grid to rejectscatter can be quantified with the scatter transmission factor t_(s),expressing how much of the scattered radiation impinging on to the gridis actually transmitted. Thus, relatively low values of the scattertransmission factor t_(s) may be desirable. In general, the scattertransmission factor t_(s) scales with the primary transmission factort_(p). In the following, the primary transmission factor t_(p) thatwould match the scatter transmission factor t_(s) of 8%, oneapproximation of the transmission factor t_(s) that an IGCT system canachieve without an anti-scatter grid, can be calculated. A relationshipbetween t_(s) and t_(p) can be calculated from the scatter-to-primaryratio without grid (SPR) and with grid (SPR_(g)):

$t_{s} = {\frac{{SPR}_{g}}{SPR} \cdot {t_{p}.}}$

For the calculation, the data of Chan et al, who used a 17-cm-thickwater phantom and measured a scatter-to-primary ratio of 3.81 withoutanti-scatter grid can be used. The use of an anti-scatter grid with aprimary transmission factor t_(p) of 50% can be reduced that to anSPR_(g) of 0.24. Thus, 6% can be derived as the corresponding value ofL.

A grid with a transmission factor of 64% can yield a significantlyhigher t_(s) of 14%. Thus, to achieve a t_(s) of 8% an anti-scatter gridwith t_(p) of 54% would have to be used. However, in CT applications ahigher scatter to primary ratio can be expected than measured by Chan etal. For example, scatter rejection can get slightly more efficient withhigher scatter-to-primary ratios. Doubling of the SPR could lead to a10-15% improvement in scatter rejection compared to that calculatedabove. Scatter rejection in IGCT may be comparable to use of ananti-scatter grid in a conventional system with a primary transmissionfactor t_(p) of 60%. Thus, intrinsic scatter rejection in IGCT canincrease the dose efficiency by 1.7-fold.

In pediatric imaging the use of an anti-scatter grid may not bebeneficial as scatter may be low. In fact, where possible it isrecommended to remove the anti-scatter grid. This may not be possible inconventional CT; particularly in pediatric imaging IGCT may be of greatadvantage to dose savings.

In the IGCT system, a photon-counting detector can be used. Compared toenergy-integrating detectors, the noise performance of photon-countingdetectors can be better. This is particularly true in a low count regimeas is often encountered in very opaque regions of the patient.Additionally, photon-counting detectors can have a bias towards lowerenergies, thus offering better contrast resolution. The stopping power,e.g. X-ray photon-stopping power, of the detector can be comparable toconventional CT detectors. Thus, fewer photons may be needed for thesame image quality and can be counted as a (virtual) flux increase of afactor of 1.2.

IGCT can allow the implementation of global and local adaptive exposure.Local adaptive exposure can vary exposure locally, e.g. depending on theopacity of the tissue scanned. Global adaptive exposure can be theoverall change of the scan pattern, e.g. depending on the global shapeof the patient. The effect can be similar to a bowtie filter used in CTimaging. In contrast to a bowtie filter, global adaptive exposure can beadjusted with the view angle.

Long path lengths, e.g. relatively long distances photons must travelthrough patient volume before reaching a detector, can lead to highattenuation of the incident beam, which can lead to low count rates. Themajority of source positions lead to count rates that are at least tentimes higher than the lowest count rate. Using adaptive exposure,acquisition times can be increased in low-count-rate areas and decreasedin high-count-rate areas, thereby equalizing the number of detectedphotons. In an ideal case, the acquisition times of most sourcepositions (90%) could be reduced to 10% of those in the low-count-rateareas. However, realistic implementation of a virtual bowtie may providesomewhat less acquisition-time savings. Acquisition-time savings of afactor of five can be achieved. However, dose savings can be much lesssince a significant fraction of the global adaptive exposure may occuroutside the patient. The combined dose savings from local and globaladaptive exposure can be larger than a factor of 2.

The scan patterns generated with adaptive exposure can preserve theoverall timing and angular sampling of the CT scan. One implementationinvolves not changing the total number of illuminated collimator holesper superview. For example, the same number of holes per superview canbe illuminated as are illuminated in a scan implemented without adaptiveexposure. Illumination patterns for interleaved superviews can providean increased flux in the central region of the collimator. The use ofdifferent illumination patterns can provide a two-dimensional adaptivefilter. Several schemes for selecting the distribution of illuminationare possible. For example, the illumination for one superview can bebased on the results of the previous superview.

The combined dose-savings effect from local and global adaptive exposurecan be at least a factor of 2, and potentially higher as the dosesavings from local adaptive exposure can be as high as 2-fold by itself.In conclusion, the combined dose savings from scatter reduction,detector technology and adaptive exposure may be at least 4-fold.

An IGCT system of one embodiment of the present invention may notproduce as many photons as modern, regular geometry point source CTscanners. As pointed out earlier, reduction in solid angle provides thesignificant design advantages in IGCT. However, reduction in solid anglecan also lead to reduction in available photons with otherwisecomparable operational parameters. The combination of scatter reduction,adaptive exposure, and a more efficient detector may allow the use ofphotons more efficiently.

The detector of a standard, e.g. regular geometry, point source CTsystem may be is 32 cm wide and 90 cm long. The IGCT system of oneembodiment of the present invention has a detector size of 6 cm by 16cm. Thus, the IGCT detector area in this embodiment may be 0.08 ofanother point source CT detector.

The duty cycle (the source-spot “on” time) of a scanning-beam digitalX-ray system may be 80% compared to 100% in a standard, e.g. regulargeometry, point source system.

A standard, e.g. regular geometry, point source CT system may have apower rating of 75 kW. The source of one embodiment of the presentinvention can operate at 25 kW. A standard point source system mayoperate at 3 rps—a rotation speed mainly geared towards cardiac CT. Withcancer imaging, rotation speeds of 1.5 rps can be sufficient. (Thenumber of photons in IGCT may be reduced by a factor of 0.67.) Scanningbeam sources can operate up to 50 kW or even 75 kW, and can operate atrotation speeds including 3 rps.

An IGCT system may have a slightly shorter focus-to-detector distancethan a regular-geometry point source system, giving the IGCT system afactor of 1.2 advantage.

As discussed, the smaller detector can lead to significantly reducedscatter in the projection images and allow for operation withoutanti-scatter grid. An increase in photon count by 1.67 can be achieved.As also discussed earlier, superior detector technology can provideequal image quality at lower photon flux. This factor can be 1.4.

The transmission anode of an IGCT system in one embodiment of thepresent invention can provide 1.7 times as many photons for the samecurrent as a traditional steep-angle reflection anode.

As discussed, adaptive exposure can enable implementation of moreefficient scan patterns. An average increase of a factor of 5 can beachieved.

The cumulative advantages and disadvantages, summarized in Table 1Error!Reference source not found., show that the number of available photonsis comparable to that of a standard, e.g. regular-geometry, point sourcesystem.

TABLE 1 IGCT system properties compared to regular- geometry pointsource system IGCT/Standard Point Source IGCT property relative toStandard System Cumulative Point Source System 0.08 0.08 smallerdetector area 0.80 0.06 lower duty cycle for IGCT 0.67 0.04 less tubepower 1.20 0.05 shorter source-detector distance 1.70 0.09 operationwithout AS grid 1.40 0.13 photon-counting detector 1.70 0.21transmission anode 5.00 1.07 adaptive exposure Equivalent 1.07Advantages offset photon loss Photon Count

The duty cycle of an IGCT system in one embodiment of the presentinvention can be increased to 100% of the duty cycle of a conventional,e.g. regular geometry, point source CT system. The IGCT system can usemultiple tubes that can be alternated thus filling in the off-time of asingle source. In addition, both iterative reconstruction and energyresolving detectors can improve performance.

Overall, other implementations of IGCT can increase the effective numberof photons by more than a factor of two compared to the numbers in Table1.

IGCT data of embodiments of the present invention can be used withiterative-reconstruction methods. In the past, these methods have notbeen used for CT. However, incorporation of fast computing platforms maymake it possible to reconstruct large datasets, e.g. IGCT datasets, inreasonable times.

Maximum Likelihood Expectation Maximization (MLEM) may be utilized toreconstruct datasets from unconventional geometries. An MLEM-basedmethod can be less prone to under-sampling artifacts and reduce noisecompared to standard methods. For example, iterative-reconstructionmethods can cut the dose in CT scans by one third;iterative-reconstruction methods may be able to create better imagesfrom less complete dataset, relative to back-projection or otherstandard methods, such that fewer detected photons may be required topopulate a dataset sufficient for reconstruction.

In summary, dose savings of at least 4-fold can be achieved.

The purpose of a collimator is to collimate X-rays emitted from thesource onto the detector. In one embodiment of the present invention, acollimator can be separated into two subsystems: a stationary collimatorring and a rotating collimator arc. An advantage is that the rotatingcollimator arc can be relatively lightweight; thus, it can reduce theweight of the subsystem that rotates at high speeds. In addition tocollimation, the collimator can serve as a containment structure forwater cooling of the X-ray target. Water cooling can be sandwichedbetween the target and a window that will be mechanically supported bythe ring collimator. The window may comprise a layer of thin aluminum,or any other thickness and material which can contain the cooling waterwhile allowing X-rays to propagate from the target screen to thecollimator. For example, the window may also be a layer of very thinstainless steel, copper, titanium, metal alloy, alloy, or any othermaterial.

The stationary collimator ring can consist of a number of slotsperpendicular to the central axis and extending over the entire lengthof the ring. The number of slots may be 10, 11, 12, 13, 14, 15, or anyother number of slots up to 50 slots or any range between 10 and 50. Forexample, in one embodiment of the present invention, the stationarycollimator ring comprises 16 slots. Each slot can be angulated anddesigned to collimate the X-rays onto the X-ray detector along the shortdimension of the detector. Collimation along the long dimension can beachieved with the rotating collimator arc. The short dimension of theX-ray detector may be the dimension parallel to a central axis throughthe stationary collimator ring, and the long dimension of the X-raydetector may be the dimension perpendicular to the central axis, e.g.around the stationary ring.

The stationary collimator ring may be fabricated of an X-ray attenuatingmaterial. The number of rings may be 6 to 201 (e.g. creating 5 to 200slots), inclusive, or any other number of rings therein or any rangebetween 6 and 201. An X-ray attenuating material may have an atomicnumber between 11 and 83, inclusive. It may be a material of an atomicnumber between 11 and 38, inclusive, for sufficient attenuation in mostcases, or a material of an atomic number between 39 and 83, inclusive,for relatively stronger attenuation or any range of atomic numbersbetween 11 and 83. Rings comprising a material of an atomic numberbetween 39 and 83 may be relatively thinner than rings comprising amaterial of an atomic number between 11 and 38. For example, lead,copper, brass, or any other attenuating material or combination ofmaterials may be used.

As discussed later, the stationary collimator may also providestabilization and support for a cooling fluid system against the X-raysource target. Collimator ring material may have a high value of Young'smodulus in order to provide this support. (The stiffness of a materialrelates to the amount of strain, e.g. the amount of deformation relativeto its original dimensions, exhibited by the material when an externalstress is applied and can be characterized by a quantity called Young'smodulus.) The Young's modulus of the material may be at least 200 GPa.Alternatively, a material with Young's modulus of 150, 150-185, 159,181, 193, 200, 190-210, 207, 248, 276, 287, 329, 345, 400-410, 435, 450,450-650, 517, 550, 1000, 1050-1200, 1220 GPa or values between 150 and1220 GPa or any range between 150 and 1220 GPa can be used. Carbonfiber, diamond, silicon carbide, steel, tungsten, tungsten carbide,iron, silicon, beryllium, molybdenum, sapphire, osmium, graphene,chromium, iridium, tantalum, or other materials can be used. Forexample, in one embodiment of the present invention, the stationarycollimator comprises 17 stainless steel rings.

Fabrication of a collimator by machining and assembling rings may becompared to prior multi-focal spot collimator fabrication methods, whichmay have comprised electrical discharge machining, chemical etching, andother more complicated processes. In addition to the additional freedomof position and in-use advantages provided by slot collimators for IGCT,they may be significantly less labor-intensive or costly to fabricate.

Some or all slots on the stationary collimator may be sized andangulated such that an X-ray beam from an underlying focal spot mayilluminate the entire face of a detector in the short dimension of thedetector. Alternatively, some slots may be narrower or more steeplyangulated such that X-ray beams from underlying focal spots illuminateonly a portion or subset of the detector in the short dimension. Theoption of illuminating portions or subset of the detector face ratherthan the entire face may be particularly useful for the implementationof adaptive exposure; for example, subsets of an image volume may beilluminated more frequently, e.g. by all available focal spots, whereasother subsets may not be illuminated when focal spots within a narrowlycollimated row are fired. It may also provide a manner of controllingthe field of view of the system, e.g. decreasing the field of view byonly firing focal spots within narrowly collimated rows.

Variations in the sizing and angulation of rows as described may followa smooth continuum or may be in any other predetermined order, forexample, “randomly” distributed in appearance.

The rotating collimator arc can extend over an arc of 60 to 160 degrees.It can also have slots, but the slots may be oriented perpendicular tothe slots of the stationary collimator ring. Thus, this collimatorcollimates the X-rays along the long dimension of the X-ray detector.The “rotating” collimator arc does not necessarily need to rotatethrough 360 degrees for a given application, though it may rotatethrough 360 degrees or more in some embodiments of the presentinvention. It may move or slide along the stationary collimator ringthough any number of degrees between 60 and 160 degrees or any rangebetween 60 and 160 degrees in embodiments of the present invention.

FIG. 9 is a diagram illustrating a stationary collimator ring androtating collimator arc of one embodiment of the present invention. InFIG. 9, stationary collimator ring 905 comprises slots subtending 360degrees around a central axis, e.g. central to the ring, whereascollimator arc 906 comprises relatively short slots aligned parallelwith a central axis. The overlap of collimator slots between stationarycollimator ring 905 and collimator arc 906 may produce a pattern ofcollimator holes, e.g. collimator hole 907, as previously described.Collimator hole 907 can illuminate detector 908.

An IGCT scan can be performed by rotating the collimator arc inside thecollimator ring. At every position of the collimator arc, thecombination of slots of the collimators can form holes that preciselycollimate the X-ray beam onto the X-ray detector. The overlap of thestationary and rotating collimators may create square holes, possibly ofside length between 0.5 mm and 1 cm. Holes may be square, rectangular,or diamond-shaped. Side lengths may include but are not limited to 1 to10 mm, 10 to 20 mm, 20 to 30 mm, 30 to 40 mm, 40 to 50 mm, 50 to 60 mm,60 to 70 mm, 70 to 80 mm, 80 to 90 mm, and 90 to 100 mm, inclusive, orany other length within the enumerate ranges or any range between 0.5 mmand 1 cm. Side lengths may also include 0.5 to 2.5 mm, 2.5 to 5.5 mm,5.5 to 7.5 mm, 7.5 to 9.5 mm, 9.5 to 11.5 mm, 11.5 to 13.5 mm, 13.5 to15.5 mm, 15.5 to 17.5 mm, 17.5 to 19.5 mm, 19.5 to 21.5 cm, inclusive,or any length within the enumerate ranges or any range between 0.5 to21.5 cm. As the collimator arc is rotating the collimator holes aremoving in time across the surface of the collimator ring.

The electron beam and the resulting X-ray beam can move quickly comparedto the collimator rotation speed. In such embodiments of the presentinvention, superviews can be generated quickly relative to collimatormotion. For example, in one embodiment the overlap of the stationarycollimator ring and rotating collimator arc can create 400collimator-hole positions, the scanning of which can generate asuperview within 0.5 ms. If the collimator arc and detector are operatedat a rotation speed of 1.5 rps, the collimator may have moved 2.3 mmduring the time, e.g. 0.5 ms, taken to complete a superview scan. Otherembodiments can create and scan 100 to 2000 hole positions. Embodimentsmay have between 100 and 200, 200 and 300, 300 and 400, 400 and 500, 500and 600, 700 and 800, 800 and 900, or 900 and 1000 holes, inclusive, orany number within the enumerated ranges or any range between 100 and1000. For example, embodiments may have 300, 350, 400, 450, 500, 550,600, or 650 holes.

To operate the scanner safely and efficiently, the X-ray beam should bealigned with the holes formed by the collimator ring and collimator arc.Alignment can be accomplished in two steps: First, the position of theX-ray beam may be established with respect to the rotating collimator atcertain time points. Second, with knowledge of the beam position and thecollimator geometry, the beam can be navigated to new collimator holes.As the collimator arc is moving, the navigation can take the collimatorvelocity into account. The velocity of the collimator arc can bemonitored using LED-photodiode sensors or similar.

To measure X-ray beam position the collimator arc can be equipped with asmall pixelated detector chip. For example, it may be equipped with apixelated detector chip that is 1 cm², as small as 1 mm², as big as 6cm², or any area in between. In some embodiments of the presentinvention the detector chip is 0.25 cm², 0.5 cm², 0.75 cm², 1 cm², 1.25cm², 1.5 cm², 1.75 cm², 2 cm², or any other size between the enumeratedvalues or any range between 1 mm² and 6 cm². It may be important todifferentiate between the measurement of the beam position at startup ofthe scanner versus during the scan. At start up, the beam can be locatedon the detector in following way: the beam can be parked on the targetat low power and highest permissible repetition rate. The collimator arccan rotate and at some point bring the detector into alignment with theX-ray beam. Once the beam is initially located, the detector can berepeatedly visited at the beginning of every scan of a superview, forexample by calculating the expected position of the detector and aimingthe X-ray beam to that location. With the detector illuminated, thecentroid of the measured beam profile during a scan can be calculated.The centroid position can give the measured beam position, which can becompared with the calculated beam position, and the calculated beamposition can be corrected for discrepancies.

This alignment scheme can be very stable as it can correct for actualbeam position at a high frequency. Loss of alignment can be quicklydetected and corrected or the scan can be aborted for patient safety.

FIG. 10 is a diagram illustrating a collimator alignment configurationof one embodiment of the present invention. In FIG. 10, parallel slots91 align with beam spots, e.g. beam spot 92, such that an X-ray beamemitted at that beam spot can pass through the collimator. (Forsimplicity, perpendicular slots from the stationary collimator are notshown, though they may be present between rows of beam spots, e.g.perpendicular to parallel slots 91.) It may be important that only abeam spot within one of parallel slots 91 be illuminated, as anun-collimated beam spot may irradiate more area than intended for theimaging procedure and possibly provide an unhealthy amount of X-rayexposure to the person being imaged and/or surrounding personnel.

Detector 90 may be attached to the rotating collimator in one ofparallel slots 91, on the edge of the rotating collimator, or in anyother location on the collimator such that it can be periodicallypositioned over a beam spot. Periodically, e.g. once per superview, abeam spot aligned with the calculated position of the detector, based onthe calculated position of the collimator, may be illuminated by thescanning beam. If the calculated position of the collimator is correct,or close to correct, detector 90 may receive some amount of radiation.

Detector 90 may be a pixelated or otherwise spatially resolved detector.FIG. 11 is a diagram illustrating a signal from a detector, e.g.detector 90, if the calculated collimator position was correct, in oneembodiment of the present invention. When a focal spot is fired under acorrectly calculated position for detector 90, centroid 93 of signal 95may align with center 94 of detector 90 whereas if the detector positionwas incorrectly calculated, it may not. FIG. 12 is a diagramillustrating a signal from a detector, e.g. detector 90, if thecalculated collimator position was incorrect, in one embodiment of thepresent invention. In FIG. 12, signal centroid 93 is not aligned withdetector center 94. However, the distance of centroid 93 from detectorcenter 94, the overall position of signal 95 on detector 90, or anyother metric may be used to recalculate or recalibrate the assumeddetector position. An interlock or safety system may be implemented suchthat X-ray emission can be shut off if centroid 93 is significantlynon-aligned with detector center 94 or if no signal is detected bydetector when the beam spot at the calculated detector position isfired.

In some imaging applications, a field of view smaller than entire widthof the scanner may be desirable. To accommodate these applications, anadditional set of slots or subset of slots that would illuminate areduced width of the detector can be used, thereby reducing the field ofview in the z-direction. As previously described, a subset or subsets ofslots may be sized or angulated in a manner to illuminate only a portionof the detector width. Alternatively, an additional stationarycollimator ring may be interchanged for or overlaid on the original.

The collimator may be constructed in a way that it sufficiently shieldsradiation not aimed at the detector. The goal of collimator design maybe to maximize efficiency, minimize leakage (radiation penetrating thecollimator) and spill (radiation through the collimator hole notcaptured by the detector). Further, the overall height of the collimatormay be sufficient to minimize penumbra at the detector. Leakage of thecollimator can be in the low-percentage range. Further reduction ofleakage can be achieved by adding sheets of high-Z (high atomic number)material such as tungsten or lead.

In addition, the collimator can be designed to attenuate photon energieswith upper limit as low as 10 keV and photon energies as high as 240 keVor any energies in between. The collimator can have 10 to 10,000 holeswith a hole pitch, e.g. center-to-center spacing, between 1 mm and 10cm. Hole pitch may be determined by the width of material strips orbeams which form slots in the stationary collimator ring and rotatingcollimator arc. Hole pitch may be between 1 mm and 5 mm, 5 mm and 10 mm,10 mm and 50 mm, 50 mm and 1 cm, 1 cm and 5 cm, 5 cm and 10 cm,inclusive, or any integer or non-integer length within these enumeratedranges or any range between 1 mm and 10 cm. For example, hole pitch maybe 3 mm, 4 mm, 5 mm, 6 mm, 7 mm, 8 mm, 9 mm, 10 mm, 11 mm, 12 mm, 13 mm,14 mm, 15 mm, 16 mm, 17 mm, 18 mm, 19 mm, or 20 mm, or any number ofmillimeters between these enumerated values. Alternatively, hole pitchmay be 0.2 cm, 0.3 cm, 0.4 cm, 0.5 cm, 0.6 cm, 0.7 cm, 0.8 cm, 0.9 cm, 1cm, 1.1 cm, 1.2 cm, 1.3 cm, 1.4 cm, or 1.5 cm, or any number ofcentimeters between these enumerate values.

The X-ray tube can run at acceleration voltages between 30 and 240 kVp.For example, maximum power in one embodiment of the present invention is25 kW. In this embodiment, a maximum exposure of 11.5 Roentgen/min hasbeen measured at 25 cm above the collimator. As discussed earlier, in anembodiment of the present invention an X-ray source can produce asufficiently high flux to support an IGCT scan at 1.5 rps and canoperate at 50 kW at 3 rps. Furthermore, the maximum high voltage can beincreased to 240 kVp, and power levels can range from 5 kW to 125 kW.

Another aspect of embodiments of the present invention is spatialresolution and focal-spot size. The X-ray tube can enable full controlover the focal-spot shape and size with sophisticated electron opticswith focus and stigmation coils. A focal-spot size can be chosen thatallows the system to produce sufficient spatial resolution forfluoroscopy systems and, at the same time, accommodate the thermal-loadrequirements of a given system. Spatial resolution requirements in CTmay be less than in fluoroscopy. The IGCT system may produce superiorspatial resolution compared to conventional, e.g. regular geometry, CTsystems. In one embodiment of the present invention, wherein a highresolution detector is utilized, spatial resolution of 2.4 mm/lp can beachieved.

The X-ray target may be the dominant source of heat in the X-ray source,and cooling by direct contact with water can enable continuous removalof the full power load applied to the target. Additionally, backscatterof electrons from the target can result in a smaller, but significantheat load to other parts of the tube vacuum bell, where heat can also beremoved by direct water cooling. The electronics that deflect and focusthe electron beam may also be water-cooled. Finally, forced ambient airmay be used to cool a ceramic insulator that may be included in tubedesign.

With the proper supply of cooling water and ambient air, the x-raysource in embodiments of the present invention can be used to operate atfull power continuously in the normal environment of a hospital orimaging center.

The overall requirements for cooling water and HVAC for a cardiacfluoroscopy system built with the scanning-beam tube and for an IGCTsystem of one embodiment of the present invention may be: HVAC, 7000BTU/hr and a water cooling supply running 20 gal/min intermittent, 30psig drop from system inlet to outlet, 70 psia max pressure at eitherinlet or outlet, and 5 C to 15 C allowable temperature range.

As mentioned, most of the energy of the accelerated electrons may bedeposited as thermal energy into the target layer of an X-ray source.Local and global heating of the target can be differentiated. Theelectron beam may heat the coating and immediate interface, e.g. windowor beryllium window, rapidly. Further away from the electron-impactzone, heating may occur more slowly as the affected volume can besignificantly larger. At distances of millimeters from theelectron-impact zone, time constants of the heating can be long and theoverall heating small. At the window-coolant interface, the problem canbe largely treated as a steady-state problem globally affecting theheating of the system.

As indicated by thermal simulations, two important parameters forthermal performance may be the peak energy intensity (J/m²) of the spotand the total power. The peak energy intensity of the spot can determinethe maximum transient temperatures in the target, which can be limitedby the melting temperature of the target materials. The total power candetermine the global heating of the target and may need to be exceededby the cooling capability of the system.

The local heat transfer can evolve in three distinct phases. First, theelectron beam can impinge on the target for 1 μs. In this phase, thetarget coating, e.g. tungsten coating, can heat rapidly with negligibleheat transfer to neighboring regions. In this phase, the thermal limitsmay largely be governed by the melting temperature of the targetmaterial, e.g. tungsten. In the second phase, heat may start spreading,decreasing the target coating temperature and increasing the temperaturein the interface region. Limitations in this phase may be limits in thethermal properties of the interface region, e.g. beryllium window. Inthe third phase, after hundreds of microseconds the system may belargely equalized, but temperature can still be elevated. If it isdesired to revisit the same focal spot in this timeframe, an offsettemperature may be taken into account. For the scanning-beam digitalX-ray system, a repetition rate of once every 300 μs can be achieved.For an IGCT source of some embodiments of the present invention, arepetition rate slower than once every 300 μs can be used. For example,in one embodiment of the present invention a repetition rate of onceevery 300 ms can be used. Alternatively, repetition rates of once every50 ms, 100 ms, 150 ms, 200 ms, 250 ms, 300 ms, 350 ms, 400 ms, 450 ms,500 ms, or any other number or fraction of milliseconds may be used orany range between 50 ms and 500 ms.

The transient heating of the target with electron-beam exposure can bestudied in great detail using Monte Carlo and/or finite-elementsimulation. Incident electrons may deposit most of their energythermally in the material, leading to significant heating of the target.Both the magnitude (power) of the electron heating as well as itsspatial distribution within the target material can be inputs to thethermal analysis. The distribution of energy in the z-direction(perpendicular to the material surface) and in the radial direction(parallel to the material surface) can be characterized.

The distribution in the z-direction can depend on the energy of theincident electrons and the target material. For example, the z-profileof the thermal deposition in tungsten for various electron energies canbe calculated using Monte Carlo simulations. An incident electron beamcan be generated, e.g. simulated, and its component electrons tracedthroughout the target. The deposited energy of every event can then betallied and histograms computed showing energy vs. depth. Electron beamsup to 140 kVp can be studied. The penetration depth and the width of thedepth profile peak may increase with incident electron energy. Themajority of energy may be deposited within a 10-μm depth.

The spot size of the electron beam may be controlled by the electronoptics of the electron gun. For some imaging applications, the spot sizemay be minimized as much as possible without exceeding the thermallimits of the target. Electron-beam spots may be described fairly wellby a Gaussian profile.

In some embodiments of the present invention, power can be 50 kW. Thespot dwell time can be 1 μs and the beam move time can be 0.25 μs.Alternatively, in the IGCT system of one embodiment of the presentinvention, the dwell time can be shortened to 0.25 μs, thus reducing thedeposited energy per spot. To recover the reduced duty cycle, the firingof adjacent source tubes can be alternated; thus, while the scanningelectron beam moves in one tube, the other tube fires and vice versa.However, in the IGCT system of another embodiment of the presentinvention, a power of 25 kW can be sufficient and the scanning-beamdigital X-ray technology can be readily applied.

The thermal energy from every focal spot can ultimately reach thewindow-coolant, e.g. beryllium-coolant, interface and contribute to theglobal heat transfer of the system. Cooling can be sufficient to removethe heat from the entire system. The bulk cooling of the X-ray tube canbe done by forced convection of cooling water that is injected in a thinlayer between the substrate, e.g. beryllium, and the collimator. Thiscooling mechanism can ensure that the tube operates at steady-statetemperature and thermal calculations can show that this coolingmechanism is sufficient to operate the X-ray tube constantly at a powerof 45 kW. With a power limit of 25 kW, the system can run continuouslyat full power for many hours. For example, initial conditioning protocolof the tube may include full-power operation for 12 hours. The coolingcapacity of forced convection can depend on the area it is applied to.Therefore, problems with the bulk cooling capacity of IGCT systems ofembodiments of the present invention are unlikely, even if operated athigher powers, as the surface area of an IGCT system can be about 10times larger than that of a scanning-beam digital X-ray tube.

For the IGCT X-ray source ring of one embodiment of the presentinvention, water cooling can be implemented that collectively cools theentire ring. Water can be injected into a thin layer between target andcollimator from one side of the ring and water can be collected at theother side of the ring.

A fixed source ring with a two-tiered collimator can be used. Completesampling can be produced and the CT data can be reconstructed. A slotcollimator can be used and/or a rotating collimator detector assemblycan be also used. In addition, rather than rotating, the collimatordetector assembly can be rocked in a way that every slot in the rotatingcollimator can be illuminated. The object can then be mounted on arotary stage. The combination of rocking the collimator and rotating theobject can enable production of a complete data set. A drawback of thisapproach may be that with a source having a flat target, it may notconform to the rotating collimator. This drawback can be overcome byusing just the central region of the source that can be closelypositioned to the rotating collimator. Different acquisition strategiescan be used and iterative-reconstruction algorithms can be used toreconstruct the data.

Scanning-beam X-ray digital systems can be used for interventionalradiology and the field of view of a cardiology scanning-beam X-raydigital system can be increased. Rather than extending the source, twodetectors can be used with a specialized collimator to illuminate bothdetectors. Two closely abutted X-ray sources can be used. Another areawhere the technology could be very beneficial is a stationarytomosynthesis system. Rather than rotating an X-ray source and detector,a large-area multi-focus X-ray tube array can be used and it can acquireimages without movement of the components. This could offer real time3-D images that can be used for image-guided procedures.

FIG. 13 is a diagram illustrating the illumination of two detectors bytwo closely abutted X-ray sources in one embodiment of the presentinvention. “Abuttable” sources 81 can function as a single, large-areamulti-focus X-ray source, or a single, large-area multi-focus X-ray tubecan be used. They can be collimated by one or two collimators toilluminate detectors 82. Alternatively, they may be collimated toilluminate two, three, four, or more detectors. Two, three, four, ormore sources can also be abutted and function as a single, large-areaX-ray source for a further increased field of view. The field of view ofthe system may encompass patient 83 as shown. Even if abuttable sourcesremain stationary or closely abutted with one another during most or allimaging procedures, it may be more cost-effective to utilize abuttablesources of embodiments of the present invention in an imaging systemrequiring significant source area rather than a single source of saidarea; fabrication of large sheets of target material can X-ray sourcescan be difficult and costly.

These are only a few examples for the vast possibilities of extendedmulti-focus X-ray sources.

An IGCT system design of embodiments of the present invention can havethe following advantages over existing high-performance conventionalmulti-slice systems: 4-fold lower dose than conventional CT, fastervolume acquisition than helical CT, whole-organ imaging with no tabletranslation and no cone-beam artifacts, and others. The scanning-beamdigital X-ray system uses a large-area scanning X-ray source to projectan X-ray beam through the patient onto a small-area, high-efficiencydetector. A high-speed computer reconstructs multi-slice tomographicimages in real time. This geometry and reconstruction can provide manysubstantial imaging and performance advantages as well asradiation-reduction advantages for the patient, physician, andfluoroscopy lab staff.

An inverse geometry CT system of embodiments of the present inventioncan use a fixed source ring and a rotating detector collimator assembly.It can also use: the combination of a stationary and a rotatingcollimator; the engineered gap to conform to Tuy's criterion; the use ofa detector on the rotating collimator for alignment; the combination ofperpendicular slots to form holes; and adaptive exposure. A compacthigh-power scanned-electron-beam X-ray sources with a large number offocal spots arranged in a two-dimensional array can be used.

Components of a scanning-beam digital X-ray source of embodiments of thepresent invention may be the cathode assembly, target assembly,collimator, shielding, cooling manifolds and mounting brackets. TheX-ray source can share the same cathode assembly, but can use asignificantly different target assembly. Collimator and cooling can beshared between the 9 sources, or other numbers of sources utilized.Mounting and shielding can be part of the overall gantry.

FIG. 14 is a diagram illustrating components of a scanning-beam digitalX-ray source of one embodiment of the present invention. Cathodeassembly 84 may be attached to the target assembly 85, comprising target87 within vacuum bell 86. As previously discussed, target 87 and thewindow of vacuum bell 86 may be curved to better fit around and alignwith a stationary collimator ring. The edges of target 87 may be flat,or may be chevron-shaped or curved, as in the embodiment of FIG. 14.Target 87 may be welded directly to vacuum bell 86 such that a verysmall amount of dead space exists on target 87, and the source can beclosely abutted with other sources.

In another embodiment of the present invention, the fixed X-ray sourcering comprises a single vacuum envelope with a continuous X-ray target.The envelope can house several electron guns that can illuminate thetarget in sequence.

The cathode assembly can consist of electron gun, acceleration anodes,gun and acceleration electronics, focus and deflection coils, focus anddeflection electronics and the enclosure. The target assembly canconsist of a vacuum bell and target. A vacuum bell may the intermediatepart between acceleration anodes and target. The vacuum bell can bemachined. The vacuum bell can also be casted.

In one embodiment of the present invention, the stationary collimatorcan have 16 slots, and it can be machined from stainless steel. Also,the collimator can be shared between the tubes. Similarly, cooling canbe shared between the tubes.

The foregoing descriptions of specific embodiments of the presentinvention have been presented for purposes of illustration anddescription. They are not intended to be exhaustive or to limit theinvention to the precise forms disclosed, and many modifications andvariations are possible in light of the above teaching. The embodimentswere chosen and described in order to best explain the principles of theinvention and its practical application, to thereby enable othersskilled in the art to best utilize the invention and various embodimentswith various modifications as are suited to the particular usecontemplated. It is intended that the scope of the invention be definedby the claims appended hereto and their equivalents.

What is claimed is:
 1. An X-ray imaging system for imaging an object comprising: a vacuum bell for creating a vacuum envelope in an X-ray source; an X-ray radiation-permeable window configured to emit X-ray radiation from a plurality of spots; a collimator located between said X-ray source and said object for projecting said X-ray radiation through said object; an X-ray detector for measuring amount of said X-ray radiation passing through said object and striking said detector; a second vacuum bell in contact with said vacuum bell for creating a second vacuum envelope in a second X-ray source; and a second X-ray radiation-permeable window configured to emit X-ray radiation from a second plurality of spots.
 2. The X-ray imaging system of claim 1 wherein said X-ray radiation-permeable window and said second X-ray radiation-permeable window are in contact with each other.
 3. The X-ray imaging system of claim 1 wherein said plurality of spots is located less than 1 cm from an edge of said window.
 4. The X-ray imaging system of claim 1 further comprising: a bonded connection between said window and said vacuum bell;
 5. The X-ray imaging system of claim 4 wherein said bonded connection is a brazed bond.
 6. The X-ray imaging system of claim 4 wherein said bonded connection is an electron beam weld.
 7. A computed tomography X-ray imaging system for imaging an object comprising: a plurality of stationary X-ray sources forming a ring for producing X-ray radiation; a rotating X-ray detector positioned within said ring for measuring amount of said X-ray radiation passing through said object and striking said detector; a stationary collimator located between said plurality of X-ray sources and said object; a rotating collimator located between said plurality of X-ray sources and said object with a plurality of slots.
 8. The computed tomography X-ray imaging system of claim 7 wherein said stationary collimator further comprises at least ten slots.
 9. The computed tomography X-ray imaging system of claim 7 wherein said stationary collimator further comprises between 10 and 50 slots.
 10. The computed tomography X-ray imaging system of claim 7 wherein said rotating slot collimator spans an arc between 60 and 160 degrees.
 11. The computed tomography X-ray imaging system of claim 7 wherein said stationary collimator further comprises metal rings.
 12. The computed tomography X-ray imaging system of claim 7 further comprising: cooling water coupled to an X-ray target for removing heat generated by said X-ray target.
 13. The computed tomography X-ray imaging system of claim 7 wherein said stationary collimator further comprises at least ten slots perpendicular to said plurality of slots.
 14. The computed tomography X-ray imaging system of claim 7 further comprising a sensor for monitoring velocity of said rotating collimator.
 15. The computed tomography X-ray imaging system of claim 7 wherein said plurality of X-ray sources can be operated at full power for at least one hour.
 16. A method of aligning an X-ray beam in an X-ray imaging system comprising: establishing a position of said X-ray beam with respect to a moving collimator at a plurality of time points; navigating said X-ray beam to a calculated position of a hole in said collimator; and correcting alignment of said X-ray beam based on location of said X-ray beam on said detector.
 17. The method of claim 16 further comprising: aborting said X-ray beam if said X-ray beam is not aligned.
 18. The method of claim 16 further comprising: determining centroid position of said X-ray beam on said detector; and comparing said centroid position with a calculated centroid position for said X-ray beam.
 19. The method of claim 16 further comprising: monitoring velocity of said collimator.
 20. The method of claim 16 further comprising: calculating position of said detector based on said velocity and an initial position of said collimator. 